1. Field of the Invention
The present invention relates to nuclear medical diagnosis apparatuses, and in particular, it relates to a nuclear medical diagnosis apparatus such as a PET apparatus and a SPECT apparatus capable of attaining reduction in size of a circuit by simplifying data processing, and at the same time, increasing sensitivity by reducing a count loss of data.
2. Description of the Related Art
A PET (positron emission tomography) apparatus detects the gamma ray emitted from the subject and reconfigures a tomogram showing an accumulated status of the drugs for PET, after or while radiopharmaceutical, which labels a material (for example, glucose, amino acid, and the like) easily accumulatable in the specific area (for example, cancer lesion) of a subject (for example, an examinee) by positron-emitting radionuclide, that is, the drug for PET to a subject, is administered. As the positron-emitting radionuclide, for example, oxygen-15 (15O), nitrogen-13 (13N), carbon-11 (11C), and fluorine-18 (18F) are used. As representative drugs for PET, 18F-fluorodeoxyglucose (18FDG) which accumulates in cancer lesion is known.
The positron-emitting radionuclide contained in the PET pharmaceuticals accumulated in the cancer lesion emits positron. This positron interacts with neighboring electron and annihilates. At this time, a pair of gamma rays (pair annihilation gamma rays) having an energy of 511 keV are emitted from the subject in a direction about 180° opposite, respectively. Consequently, two gamma rays each having an energy of about 511 keV detected approximately at the same time are highly probable to be a pair annihilation gamma rays generated by a single event (pair annihilation of positron and electron). Consequently, based on the position of two radiation detectors (a pair of detectors) separately detecting the two gamma rays that meet these conditions (synchronicity and energy), each track of these gamma rays can be presumed.
Likewise, by collecting the track information on a large number of pair of gamma rays, and based on these pieces of the track information, if the image reconfiguration represented by a filtered back projection (FBP) method is performed, a tomogram representing an internal radiation concentration distribution caused by the positron-emitting radionuclide can be obtained.
To generate a good PET image, it is necessary to specify a pair of detection data corresponding to the pair annihilation of gamma rays for every event. Hence, in the PET apparatus, a coincidence counting circuit specifies a pair of detection data corresponding to the pair of gamma rays practically and simultaneously detected. In this way, the pair annihilation of gamma rays is accurately recognized and used for the generation of the tomogram, and even when they are the diffused gamma ray or the annihilation gamma ray, the detection data which has detected only either one is removed.
When the coincidence counting circuit receives two detection data having time information within a predetermined time window, the coincidence circuit performs a coincidence counting with the detection data taken as received practically at the same time. The time window, for example, is a width of 10 “ns”, and is set up as short as possible to avoid accidental coincidence counting in consideration of the tracking time difference between two gamma ray of the pair annihilation of gamma rays, a limitation of the time accuracy of the signal processing system of the apparatus, and the like.
The accidental coincidence counting means that since a plurality of events (for example, an emission of gamma rays) of the same type has happened at the same time, the observational result caused by another event is taken as the observational result caused by the single event, and is erroneously recognized. For example, when two positrons annihilate at the same time in the body and the gamma rays caused by the annihilation of these positrons are detected one by one, a problem arises that it is difficult to judge that this phenomenon is attributable to the accidental coincidence counting.
A SPECT (Single Photon Emission Computed Tomography) examination is an examination that administers an radioactive drugs (drugs for SPECT), which labels a material easily accumulatable in the specific area in a live body by a single photon emission nuclide, to the examinee, and after that (or while administering), detects gamma ray emitted from the examinee, thereby reconfiguring a tomogram showing a collection and distribution status of the drugs for SPECT.
The single photon emission nuclide is broken down with an intrinsic probability by generating an electron capture (EC) and the like, and emits a single photon of the gamma ray. This nuclide includes technetium-99m (99mTc), gallium-68 (68Ga), thallium-201 (201Tl), and the like. The half-life periods of these nuclides are generally longer than the half-life period of the positron emitting radionuclide used for the PET examination, and for example, are 6.0 days (in case of 99mTc), 3.3 days (in case of 67Ga) or 73 days (in case of 201Tl), and the like. In the SPECT examination, by providing a collimator for the radiation detector and limiting an incident angle of the gamma ray, the track of the gamma ray is presumed. These single photon emitting nuclides emit the gamma ray having energy of 100 keV order.
The nuclear medical diagnosis apparatus such as the PET apparatus, for example, includes detector units to the extent of 30 units to 100 units (see JP-A-2005-106644). This detector unit packs radiation detectors for every predetermined number for about every several hundreds to several thousands.
The number X of coincidence counting circuits necessary in principle for confirming a combination of all the detector units can be determined by X=NC2 provided that the number of detector units is taken as N. Consequently, for example, if the detector units provided for the nuclear medical diagnosis apparatus are 100 units, the calculation result of this number X is about 5000. However, in the actual nuclear medical diagnosis apparatus, the coincidence counting circuits need only be about half this number. This is because, due to the geometrical relative position of the two detector units, there are a considerable number of combinations in which the segment connecting these detector units is unable to pass through a subject.
Heretofore, the coincidence between has been performed by using an analogue circuit. In this method, while a circuit scale need only be small, there are a lot of fluctuations in time, and the adjustment thereof has been difficult. Hence, a method of performing the coincidence between by digital circuit has come into practical use. According to this method, based on the timing when a radiation detection signal is received, the detection time is digital-converted to generate detection time data, and by comparing the generated detection time data with each other, the coincidence between is performed. According to this method, the width of the time window relative to the coincidence counting can be easily set up, so that the coincidence counting of higher accuracy can be performed. However, according to this method, the circuit scale of the coincidence detection circuit becomes vast.
Hence, U.S. Pat. No. 5,241,181 specification discloses a Coincidence Detector for a PET Scanner in which, heretofore, the digitalized time signal from each detector unit has been stored in a shift register, and all the combinations have been compared by each comparator circuit, thereby performing the coincidence between.
According to the conventional “Coincidence Detector for a PET Scanner”, the signal from each unit is coincidence-judged by a time sharing to attain the reduction of the number of circuits, and the time signal data from each detector unit is stored in a shift register, and the comparison of all the combinations restricted by the position is performed.
However, this “Coincidence Detector for a PET Scanner” can process only one event within a time frame. Hence, to increase the number of units, it is necessary to make the time frame short in order to process a vast amount of data or reduce the number of radiation detectors stored in one unit and reduce the number of events per each unit. However, when the time frame is made short, a rate of the set of data crossing over the time frame increases, and a rate of the data abandoned without being used for the image formation also increases. That is, a count loss of the data increases, and the sensitivity of the apparatus is lowered. When the number of radiation detectors stored in one unit is reduced, it is necessary to increase the number of units in order to obtain the same performance, and therefore, the required coincidence detection circuit is also increased, thereby increasing the circuit scale.
In recent years, to improve the resolution, a degree of integration of the radiation detectors to be disposed is apt to become high. In the apparatus in which the degree of integration of the radiation detectors is made high, a probability is increased that a piece of the radiation ray is diffused, and is detected as a scattered radiation by a plurality of radiation detectors. Thus, heretofore, a technique has been desired in which, by utilizing the data of the scattered radiation which is abandoned as the energy is less than the predetermined value, the sensitivity of the apparatus is improved.